Implantable biomaterials having functional surfaces

ABSTRACT

Implantable materials having defined patterns of affinity regions for binding endothelial cells and providing for directed endothelial cell migration across the surface of the material. The affinity regions include photochemically altered regions of a material surface and physical members patterned on the material surface that exhibit a greater affinity for endothelial cell binding and migration than the remaining regions of the material surface.

CROSS-REFERENCE TO RELATED INVENTIONS

This application is a continuation application of U.S. Ser. No.16/047,297, filed on Jul. 27, 2018; which is a divisional application ofU.S. patent application Ser. No. 15/219,133, filed Jul. 25, 2016, whichwill issue as U.S. Pat. No. 10,034,967 on Jul. 31, 2018; which is acontinuation of U.S. patent application Ser. No. 14/256,674, filed Apr.18, 2014, issued as U.S. Pat. No. 9,399,087 on Jul. 26, 2016; which is adivisional of U.S. patent application Ser. No. 13/168,897, filed Jun.24, 2011, now U.S. Pat. No. 8,709,066; which is a continuation of U.S.patent application Ser. No. 11/091,669, filed Mar. 28, 2005, now U.S.Pat. No. 8,147,859; which is a continuation of International ApplicationNo. PCT/US2003/030383 filed Sep. 26, 2003 (International Publication No.WO 2004/028347), which claims priority to U.S. Provisional ApplicationSer. No. 60/414,031 filed Sep. 26, 2002, all herein incorporated byreference in their entireties.

BACKGROUND OF THE INVENTION

The present invention relates generally to implantable medical devicesand more particularly to controlling surface properties of implantablebiocompatible materials suitable for fabrication of implantable medicaldevices. Implantable medical devices are fabricated of materials thatare sub-optimal in terms of the biological response they elicit in vivo.Many conventional materials used to fabricate implantable devices, suchas titanium, polytetrafluoroethylene, silicone, carbon fiber andpolyester, are used because of their strength and physiologically inertcharacteristics. However, tissue integration onto these materials istypically slow and inadequate. Certain materials, such as silicone andpolyester, elicit a significant inflammatory, foreign body response thatdrives fibrous encapsulation of the synthetic material. The fibrousencapsulation may have significant adverse effects on the implant.Moreover, conventional biomaterials have proved inadequate in elicitinga sufficient healing response necessary for complete device integrationinto the body. For example, in devices that contact blood, such asstents and vascular grafts, attempts to modify such devices to promoteendothelial cell adhesion may have a concomitant effect of making thedevices more thrombogenic.

When implanted, conventional blood-contacting implantable devices, suchas stents, stent-grafts, grafts, valves, shunts and patches, fail todevelop a complete endothelial layer, thereby exposing the devicematerial to thrombus formation or smooth muscle cell proliferation, andultimate failure of the implanted device. It has been recognized that,when implanted into the body, metals are generally considered to havesuperior biocompatibility than polymers used to fabricate commerciallyavailable polymeric grafts.

In investigating cellular interactions with prosthetic materialsurfaces, it has been found that cell adhesion to the material surfaceis mediated by integrins present on cell membranes that interact withthe prosthetic surface. Integrins are the most prominent member of aclass of extracellular matrix (ECM) adhesion receptors. Integrins are alarge family of heterodimeric transmembrane proteins with different αand β subunits. Integrins are regulated at several levels. Modulation ofthe affinity of the adhesion receptor for ligand, termed affinitymodulation, is a mechanism for activation of platelet aggregation and isbelieved to underlie activation of leukocyte adhesion. Adhesivestrengthening by clustering of adhesion receptors or bycytoskeletal-dependent processes such as cell spreading has been shownto be crucial for strong cellular attachment, control of cell growth andcell motility. Under high shear forces present in flowing blood,leukocytes first tether, then roll along the vessel surface. When alocal signal, e.g., a cytokine, is released in their vicinity, theleukocyte arrests, develops a firm adhesion then migrates across theendothelium. Tethering, rolling, arrest and adhesion tightening are allknown to result from activation of leukocyte integrins.

Once adhered to a surface, cell spreading and migration are associatedwith assembly of focal adhesion junctions. Cell migration entails thecoordination of cytoskeletal-mediated process extension, i.e., filopodiaand lamellopodia, formation of adhesive contacts at the leading edge ofa cell, breaking adhesive contacts, and cytoskeletal retraction at thetrailing edge of the cell. Focal adhesions are comprised of integrins asthe major adhesion receptors along with associated cytoplasmic plaqueproteins. Assembly of focal adhesions is regulated by extracellularligand binding events and by intracellular signaling events. Ligandbinding controls localization of β1- and β3-containing integrins intofocal adhesions. The cytoplasmic domains of the β subunits haveintrinsic signals for focal adhesion localization, but incorporation ofthe integrins into focal adhesions is prevented by the a subunits of theheterodimers. Ligand binding, however, relieves this inhibition andallows the subunit cytoplasmic tail signals to recruit the integrindimmer into the focal adhesion.

Attempts at coating implanted metal devices, such as stents, withproteins that contain the Arg-Gly-Asp (RGD) attachment site have beenmade with some success. The RGD sequence is the cell attachment site ofa large number of adhesive extracellular matrix, blood, and cell surfaceproteins and many of the known integrins recognize the RGD sequence intheir adhesion protein ligands. Integrin-binding activity may also bereproduced by synthetic peptides containing the RGD sequence. However,bare metal implanted materials will not, of course, have native RGDattachment sites. Thus, metal implantable devices, such as stents, havebeen derivitized with polymers having RGD attachment sites bound to thepolymer matrix.

It has been found that when prosthetic materials are implanted, integrinreceptors on cell surfaces interact with the prosthetic surface. Whencells come into contact with the extracellular matrix, such as aprosthetic surface, their usual response is to extend filopodia, andintegrins at the tip of the filopodia bind to the extracellular matrixand initiate the formation of focal adhesions. Actin-rich lamellipodiaare generated, often between filopodia, as the cell spreads on theextracellular matrix. Fully developed focal adhesions and associatedactin stress fibers ensue. These same evens occur during cell migrationas cells extend lamellipodia and form focal adhesions to derive thetraction necessary for movement. Giancotti, F. G., et al. Science,285:13 Aug. 1999, 1028-1032.

The integrin receptors are specific for certain ligands in vivo. If aspecific protein is adsorbed on a prosthetic surface and the ligandexposed, cellular binding to the prosthetic surface may occur byintegrin-ligand docking. It has also been observed that proteins bind tometals in a more permanent fashion than they do to polymers, therebyproviding a more stable adhesive surface. The conformation of proteinscoupled to surfaces of most medical metals and alloys appears to exposegreater numbers of ligands and attract endothelial cells having surfaceintegrin clusters to the metal or alloy surface, preferentially overleukocytes.

Because of their greater adhesive surface profiles, metals are alsosusceptible to short-term platelet activity and/or thrombogenicity.These deleterious properties may be offset by administration ofpharmacologically active antithrombogenic agents in routine use today.Surface thrombogenicity usually disappears 1-3 weeks after initialexposure. Antithrombotic coverage is routinely provided during thisperiod of time for coronary stenting. In non-vascular applications suchas musculoskeletal and dental, metals have also greater tissuecompatibility than polymers because of similar molecular considerations.The best article to demonstrate the fact that all polymers are inferiorto metals is van der Giessen, W J. et al. Marked inflammatory sequelaeto implantation of biodegradable and non-biodegradable polymers inporcine coronary arteries, Circulation, 1996:94(7):1690-7.

Normally, endothelial cells (EC) migrate and proliferate to coverdenuded areas until confluence is achieved. Migration, quantitativelymore important than proliferation, proceeds under normal blood flowroughly at a rate of 25 μm/hr or about 2.5 times the diameter of an EC,which is nominally 10 μm. EC migrate by a rolling motion of the cellmembrane, coordinated by a complex system of intracellular filamentsattached to clusters of cell membrane integrin receptors, specificallyfocal contact points. The integrins within the focal contact sites areexpressed according to complex signaling mechanisms and eventuallycouple to specific amino acid sequences in substrate adhesion molecules.An EC has roughly 16-22% of its cell surface represented by integrinclusters. Davies, P. F., Robotewskyi A., Griem M. L. Endothelial celladhesion in real time. J. Clin. Invest.1993; 91:2640-2652, Davies, P.F., Robotewski, A., Griem, M. L., Qualitiative studies of endothelialcell adhesion, J. Clin. Invest.1994; 93:2031-2038. This is a dynamicprocess, which involves more than 50% remodeling in 30 minutes. Thefocal adhesion contacts vary in size and distribution, but 80% of themmeasure less than 6 μm², with the majority of them being about 1 μm²,and tend to elongate in the direction of flow and concentrate at leadingedges of the cell. Although the process of recognition and signaling todetermine specific attachment receptor response to attachment sites isnot completely understood, availability of attachment sites willfavorably influence attachment and migration. It is known that materialscommonly used as medical grafts, such as polymers, do not become coveredwith EC and therefore do not heal after they are placed in the arteries.It is therefore an object of this invention to replace polymer graftswith metal grafts that can potentially become covered with EC and canheal completely. Furthermore, heterogeneities of materials in contactwith blood flow are preferably controlled by using vacuum depositedmaterials.

There have been numerous attempts to increase endothelialization ofimplanted medical devices such as stents, including covering the stentwith a polymeric material (U.S. Pat. No. 5,897,911), imparting adiamond-like carbon coating onto the stent (U.S. Pat. No. 5,725,573),covalently binding hydrophobic moieties to a heparin molecule (U.S. Pat.No. 5,955,588), coating a stent with a layer of blue to black zirconiumoxide or zirconium nitride (U.S. Pat. No. 5,649,951), coating a stentwith a layer of turbostratic carbon (U.S. Pat. No. 5,387,247), coatingthe tissue-contacting surface of a stent with a thin layer of a Group VBmetal (U.S. Pat. No. 5,607,463), imparting a porous coating of titaniumor of a titanium alloy, such as Ti—Nb—Zr alloy, onto the surface of astent (U.S. Pat. No. 5,690,670), coating the stent, under ultrasonicconditions, with a synthetic or biological, active or inactive agent,such as heparin, endothelium derived growth factor, vascular growthfactors, silicone, polyurethane, or polytetrafluoroethylene (U.S. Pat.No. 5,891,507), coating a stent with a silane compound with vinylfunctionality, then forming a graft polymer by polymerization with thevinyl groups of the silane compound (U.S. Pat. No. 5,782,908), graftingmonomers, oligomers or polymers onto the surface of a stent usinginfrared radiation, microwave radiation or high voltage polymerizationto impart the property of the monomer, oligomer or polymer to the stent(U.S. Pat. No. 5,932,299). However, all these approaches do not addressthe lack of endothelialization of polymer grafts.

In accordance with the present invention, the capacity for completeendothelialization of conventional implantable materials, includingmetals and polymers, may be enhanced by imparting a pattern ofchemically and/or physiochemically active features onto a bloodcontacting surface of the implantable material. The inventiveimplantable metal devices may be fabricated of polymers, pre-existingconventional wrought metallic materials, such as stainless steel ornitinol hypotubes, or may be fabricated by thin film vacuum depositiontechniques. In accordance with the present invention, it is preferableto fabricate the inventive implantable materials and resulting devicesby vacuum deposition of either or both of the base implant material andthe chemically and/or physiochemically active features. Vacuumdeposition permits greater control over many material characteristicsand properties of the resulting material and formed device. For example,vacuum deposition permits control over grain size, grain phase, grainmaterial composition, bulk material composition, surface topography,mechanical properties, such as transition temperatures in the case of ashape memory alloy. Moreover, vacuum deposition processes will permitcreation of devices with greater material purity without theintroduction of large quantities of contaminants that adversely affectthe material and, therefore, the mechanical and/or biological propertiesof the implanted device. Vacuum deposition techniques also lendthemselves to fabrication of more complex devices than those that aremanufactured by conventional cold-working techniques. For example,multi-layer structures, complex geometrical configurations, extremelyfine control over material tolerances, such as thickness or surfaceuniformity, are all advantages of vacuum deposition processing.

In vacuum deposition technologies, materials are formed directly in thedesired geometry, e.g., planar, tubular, etc. The common principle ofvacuum deposition processes is to take a material in a minimallyprocessed form, such as pellets or thick foils, known as the sourcematerial and atomize them. Atomization may be carried out using heat, asis the case in physical vapor deposition, or using the effect ofcollisional processes, as in the case of sputter deposition, forexample. In some forms of deposition a process such as laser ablation,which creates microparticles that typically consist of one or moreatoms, may replace atomization; the number of atoms per particle may bein the thousands or more. The atoms or particles of the source materialare then deposited on a substrate or mandrel to directly form thedesired object. In other deposition methodologies, chemical reactionsbetween ambient gas introduced into the vacuum chamber, i.e., the gassource, and the deposited atoms and/or particles are part of thedeposition process. The deposited material includes compound speciesthat are formed due to the reaction of the solid source and the gassource, such as in the case of chemical vapor deposition. In most cases,the deposited material is then either partially or completely removedfrom the substrate, to form the desired product.

A first advantage of vacuum deposition processing is that vacuumdeposition of the metallic and/or pseudometallic films permits tightprocess control and films may be deposited that have a regular,homogeneous atomic and molecular pattern of distribution along theirfluid-contacting surfaces. This avoids the marked variations in surfacecomposition, creating predictable oxidation and organic adsorptionpatterns and has predictable interactions with water, electrolytes,proteins and cells. In particular, EC migration is supported by ahomogeneous distribution of binding domains that serve as natural orimplanted cell attachment sites in order to promote unimpeded migrationand attachment.

Secondly, in addition to materials and devices that are made of a singlemetal or metal alloy layer, the inventive grafts may be comprised of alayer of biocompatible material or of a plurality of layers ofbiocompatible materials formed upon one another into a self-supportingmultilayer structure because multilayer structures are generally knownto increase the mechanical strength of sheet materials, or to providespecial qualities by including layers that have special properties suchas superelasticity, shape memory, radio-opacity, corrosion resistanceetc. A special advantage of vacuum deposition technologies is that it ispossible to deposit layered materials and thus films possessingexceptional qualities may be produced (cf., H. Holleck, V. Schier:Multilayer PVD coatings for wear protection, Surface and CoatingsTechnology, Vol. 76-77 (1995) pp. 328-336). Layered materials, such assuperstructures or multilayers, are commonly deposited to take advantageof some chemical, electronic, or optical property of the material as acoating; a common example is an antireflective coating on an opticallens. Multilayers are also used in the field of thin film fabrication toincrease the mechanical properties of the thin film, specificallyhardness and toughness.

Thirdly, the design possibilities for possible configurations andapplications of the inventive graft are greatly realized by employingvacuum deposition technologies. Specifically, vacuum deposition is anadditive technique that lends itself toward fabrication of substantiallyuniformly thin materials with potentially complex three dimensionalgeometries and structures that cannot be cost-effectively achieved, orin some cases achieved at all, by employing conventional wroughtfabrication techniques. Conventional wrought metal fabricationtechniques may entail smelting, hot working, cold working, heattreatment, high temperature annealing, precipitation annealing,grinding, ablation, wet etching, dry etching, cutting and welding. Allof these processing steps have disadvantages including contamination,material property degradation, ultimate achievable configurations,dimensions and tolerances, biocompatibility and cost. For exampleconventional wrought processes are not suitable for fabricating tubeshaving diameters greater than about 20 mm, nor are such processessuitable for fabricating materials having wall thicknesses down to about1 μm with sub-μm tolerances.

Overall rate to reach confluence for the endothelial cells on the bloodcontact surface of implanted medical device is mainly determined by twofactors, the rate of cell movement and rate of cell proliferation, withthe first being more important. The rate of cell movement furthercomprises three interrelated steps. Initially, a cell forms lamellipodiaand flopodia that protrude outward. This step involves reassembly ofactins in the forefront of lambaepolia. After protrusion of lamellipodiafrom one or multiple points from the cell membrane, the front end of thelamellipodia will form a close attachment, called focal adhesion point,to the substratum through the interaction of integrin on the cellmembrane and extrcellular matrix binding site. The final step of cellmovement involves the contraction of the posterior end through theaction of myosin II. The formation of a focal adhesion point is criticalfor the cell movement because the protruding lamellipodia will otherwisefold back. Without the tension force from the focal adhesion point, acell loses the contraction from the posterior end and hence stopsmoving.

Availability of attachment sites on the substratum is not only importantfor the focal adhesion point formation, but also important forpropagation. It has been shown that cells are forced to spread, survivebetter and proliferate faster than cells that are confined to the sameamount of surface area (Science 276:1425-1428, 1997). This may explainwhy spreading of neighbor cells stimulate a cell to proliferate, aftercells are lost from epithelium.

The formation of extracellular matrix (ECM) is, to much extent,determined by the cells within it because molecules which form ECM aresecreted by the cells. Subsequently, the structure of the ECM, and hencethe distribution of attachment sites on the ECM for the integrinbinding, determines the focal adhesion point formation, the criticalstep in cell movement. Therefore, proper distribution of integrinbinding sites on the surface of an implanted medical devicesubstantially determines the speed of re-endothelialization from theends surrounding the device.

There still remains a need for a medical device that stimulatesendothelial proliferation and movement when implanted in order to forman endothelial layer over the medical device. Furthermore, there is aremaining need for a method of fabricating such a medical device.

SUMMARY OF THE INVENTION

In accordance with an aspect of the present invention, there is providedan implantable material having at least one blood contact surfacecomprising an evenly distributed geometric feature for cell attachment.The evenly distributed feature on the blood contact surface of themedical device includes: circle dots, square dots, rectangular dots,triangle dots, parallel lines and intersecting lines, or any combinationthereof. Additionally, another aspect of the present invention providesmethods of making a device that has evenly distributed geometricfeatures on the blood contact surface.

BRIEF DESCRIPTION OF THE FIGURES

FIG. 1 is a perspective view of an embodiment of the present inventionincluding evenly distributed elevated geometric features on the surfaceof an implantable material.

FIG. 2 is cross-sectional view of FIG. 1 along line 2-2.

FIG. 3 is a perspective view of an embodiment of the present inventionincluding evenly distributed chemically defined geometric features onthe surface of an implantable material.

FIG. 4 is a cross-sectional view of FIG. 3 along line 4-4.

FIG. 5 is a photomicrograph showing an embodiment of the presentinvention including geometric features as carbon coated silicon.

FIGS. 6a-6c are photomicrographs showing cellular migration on thesurface with no inventive geometric features versus on the surface withinventive features.

FIG. 7 is a photomicrograph showing the stained focal adhesion pointsclose to the geometric features.

FIGS. 8a-8b are photomicrographs showing the formation of multiple focaladhesion points of a migrating cell and its attachment to the inventivegeometric features.

FIGS. 9a-9d are cross-sectional diagrammatic views of an embodiment ofthe present invention, the combination of a-d representing the steps tomake an inventive implantable material with elevated geometric features.

FIGS. 10a-10d are cross-sectional diagrammatic views of an embodiment ofthe present invention, the combination of a-d representing the steps tomake an inventive implantable material with chemically defined geometricfeatures.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

The present inventions takes advantage of the discovered relationshipbetween chemically or physiochemically-active geometric features definedand distributed on a blood contact surface enhanced endothelial cellbinding, proliferation and migration over the blood contact surface ofthe implantable material. The present invention involves focal adhesionpoint formation during cellular movement and the well-establishedobservation known as anchorage dependence, that spreading cellsproliferate faster than non-spreading cells. It has been found theaddition of a patterned array of ultra-thin features having ahydrophobic, hydrophilic or surface energy difference relative to thesurface onto which the ultra-thin features are added, enhances thebinding, proliferation and migration of endothelial cells to and betweenthe features and across the surface. Use of the term “ultra-thin” isintended to include material thicknesses between about 0.1 μm and 3 μm.It has been found that below about 3 μm the interactions betweenendothelial cells and the ultra-thin features is primarily chemical andelectrochemical. Features having thicknesses greater than 3 μm and up toabout 20 μm may also be employed in the present invention, it beingunderstood that as the thickness of the feature increases there is adecreasing chemical and/or electrochemical interaction between thefeature and the endothelial cells and an increasing physicalinteraction.

Additionally, it has been found that by employing UV irradiation tooxidized titanium or titanium-alloy surfaces, photochemical alterationof the surface titanium oxides alter the hydrophobicity of the exposedtitanium oxides and act as affinity binding and migration sites forendothelial cell attachment and proliferation across a titanium ortitanium-alloy surface. Where UV irradiation is employed, the thicknessof the photochemically altered regions of titanium oxide are, for allpractical purposes, 0 μm. Thus, within the context of the presentapplication, the term “geometric features” is intended to include bothphysical members and photochemically-altered regions having thicknesseshaving thicknesses down to 0 μm.

In FIG. 1, a portion of an implantable material 10 showing the surfacematerial 12 with described elevated geometric features 14 isillustrated. The geometric features are elevated from the surface of theimplantable material to a height ranging from about sub-micron to about20 μm. Preferably, the height of the geometric feature 14 ranges fromabout sub-micron to about 3 μm. The shape of geometric features can beeither circular, square, rectangle, triangle, parallel lines, straightor curvilinear lines or any combination thereof. Each of the geometricfeatures is preferably from about 10 μm to about 75 μm, and preferablyfrom about 15 μm to 50 μm in feature width 16, or feature diameter ifthe geometric feature is circular. A gap distance 18 between each of thegeometric features should generally be the same as the feature width 16,i.e., between about 10 μm to about 75 μm edge-to-edge.

FIG. 2 is a cross-sectional view along line 2-2 in FIG. 1. One of theelevated geometric features 14 is shown on the surface 12 of theimplantable material.

In FIG. 3, a titanium or titanium-alloy material 20 is heating tooxidize and form titanium dioxide on the surface of the material 20,then features 24 are formed by exposing the material 20 to UV through apattern mask. UV irradiation alters the titanium oxides in the areas offeatures 24, thereby chemically altering the geometric features 24relative to the surrounding the surrounding surface area 22 of material20. The shape of geometric features can be circular, square, rectangle,triangle, parallel lines, intersecting lines or any combination. Each ofthe geometric features is from about 10 μm to about 75 μm, andpreferably from about 15 μm to 50 μm in feature width 16, or featurediameter if the geometric feature is circular. The gap distance 28between each component of the geometric features is in the samemagnitude as the feature width 26.

FIG. 4 is a cross-sectional view of FIG. 3 along line 4-4. The describedgeometric features 24 are indicated by the dotted lines, which indicatesthat the geometric features 24 are at the same level of the surroundingsurface 22.

FIG. 5 shows geometric features that are evenly distributed across theat least one surface of the implantable material that contacts bodyfluid, preferably blood. As disclosed in FIG. 1 and FIG. 2, thegeometric features are elevated from the rest of the surface to a heightranging from about sub-micron to about 20 micrometer. Preferably, theheight of the geometric feature ranges from about sub-micron to about 3micrometer. The shape of the geometric features is not confined withinthe shape that is shown. The shape of the chemically defined domain canalso be any of circle, square, rectangle, triangle, parallel lines,intersecting lines or any combination of the above.

FIG. 6A shows the cell 32 spreading on the surface of hydrophilictreated Si. FIG. 6B shows the cell 32 spreading on the surface ofhydrophilic treated Si with circular dots that are 15 microns indiameter. Cells in FIG. 6B appear to have much more focal adhesionpoints 36 than those in FIG. 6A. Because these geometric featuresprovide for cell attachment, acting as affinity domains, the size ofeach of these affinity domains relative to the size of an endothelialcell determines the availability of affinity domains to the subsequentround of cell movement. According to the present invention, thepreferred size of each of the individual component of the geometricfeatures is about 10 μm to about 75 μm, and preferably from about 15 μmto 50 μm in feature width, or diameter if the geometric feature iscircular. As described in the background section, focal adhesion pointformation is the critical step in cell movement and cell proliferation,therefore, geometric features such as carbon dots on the hydrophilic Sisurface promote cell movement. It is known to the person skilled in theart that spreading of cells promotes cell proliferation. Promoting cellmovement and cell proliferation ultimately accelerates covering of theimplanted implantable material with endothelial cells on exposedsurfaces having the geometric features. Although the geometric featuresshown in FIG. 6B are circular, the shape of the geometric features arenot limited to this particular embodiment.

FIG. 6C is a magnification of a portion of the image of FIG. 6B.Multiple focal adhesion points 36 are again shown. Wide spreading of thecell is primarily due to the formation of multiple focal adhesion pointson the circular geometric features. Extensive spreading of the cells isbeneficial towards endothelialization because it promotes cell movementand cell proliferation.

FIG. 7 shows the stained focal adhesion points 36 of human aoticendothelial cells (HAEC) on the surface of an implantable material withgeometric features 14 that are in the form of carbon dots. The focaladhesion points are located at or very close to the geometric features14. As described in the background section, these focal adhesion pointsserve as tension points for the cell to contract from the opposite endof the cell and hence promote cell movement.

FIG. 8A shows the wide spreading of cells 32 and focal multiple focaladhesion points 36 on the surface of an implantable material withgeometric features that are in the form of NiTi dots of 25 micrometersin diameter. The NiTi dots are invisible due to the weak contrastbetween the NiTi dots and surrounding Si surface.

FIG. 8B shows a magnified slide of a human aotic epithelial cell 32, asshown in FIG. 8A. Multiple focal adhesion points 36 are shown toencapsulate the NiTi dots patterned on the hydrophilic Si surface.

Referring to FIG. 9A, a portion of an implantable material 46 withsurface 42 and 44 is shown.

Referring to FIG. 9B, according to the present invention, a machinedmask 48 having laser-cut holes 40 of defined size ranging from about 10μm to about 75 μm, and preferably from about 15 μm to 50 μm, patternedthroughout coats at least one surface 42 of the implantable material 46and is tightly adhered to the covered surface 42.

Referring to FIG. 9C, a thin film of material 14 was deposited into thespace as defined by the holes 40, as seen in FIG. 9B, in the mask 48 bythin film deposition procedures.

Referring to FIG. 9D, after deposition, the mask is removed to revealthe geometric features 49 patterned across the at least one surface 42of the implantable material 46.

As described above, the shape of the holes in the mask could be in anyof the shapes described for the geometric features including: circle,square, rectangle, triangle, parallel lines and intersecting lines, orany combination thereof. In the thin film deposition embodiment of themanufacturing the geometric features, the geometric features areelevated from the surface of the implantable material. The thickness ofthe geometric features is based upon the thickness of the holes in themask, the thickness ranging from about sub-micron to about 20micrometer. Preferably, the thickness of the holes in the mask rangefrom about sub-micron to about 3 micrometer.

In accordance with an alternate embodiment of the present invention, thesubstrate for the implantable medical device is formed of titanium,nickel-titanium alloy or other titanium-rich alloy metals, which isoxidized to convert surface titanium to titanium dioxide, then coveredwith a pattern-mask and exposed to high intensity UV irradiation. It iswell-known that titanium dioxide (TiO₂) absorbs UV radiation and hasbeen used in a variety of applications as a UV inhibitor to prevent UVtransmission across a TiO₂ barrier layer. It has been discovered thatupon exposure to UV irradiation, an originally hydrophobic andoleophilic titanium oxide layer becomes amphiphilic. The effect of UVirradiation on a titanium oxide surface is believed to occur because ofunsymmetrial cleavage of the Ti—O bond to leave Ti′ ions on the surfacein some regions. Presently, these amphiphilic surfaces are being used ina range of technological applications, such as self-cleaning paints andanti-misting glasses. It has been recognized that these amphiphilictitanium oxide layers have use in medical applications. Zarbakhsh, A.,Characterization of photon-controlled titanium oxide surfaces, ISISExperimental Report, Rutherford Appelton Laboratory, May 16, 2000 (whichmay be found on the internet at:www.isis.rlac.uk/isis2001/reports/11144.pdf).

It has been recognized by the present inventors that the amphiphilicstate of the UV irradiated titanium oxide may be advantageously employedas an alternative to depositing patterned features onto the implantablesubstrate surface. An implantable substrate fabricated of titanium or atitanium alloy is masked with a pattern mask having a plurality ofopenings passing there through. As with the above-described embodiment,the plurality of openings preferably have a size and special arrayselected to define affinity binding domains and cellular migration citesfor promoting endothelial cell binding and proliferation across thesubstrate surface. The open surface area of each of the plurality ofopenings in the pattern mask is preferably in the range of between about10 to 75 μm, and with adjacent pairs of openings being in a spaced apartrelationship such that a distance of about 10 to about 75 μm existsbetween the openings, the inter-opening distance corresponding to thesize of the opening. By interposing the pattern mask between a UV sourceand the substrate surface, a pattern of UV irradiated regions isimparted to the substrate surface, thereby altering the titaniumdioxides present at the irradiated regions and forming affinity domainsat the substrate surface.

Referring to FIG. 10A, a portion of an implantable material 56 made oftitanium or a titanium-alloy is shown having at least one surface 52 and54 that is oxidized by heating or an equivalent known by the personskilled in the art.

Referring to FIG. 10B, according to the present invention, a machinedmask 48 that had laser-cut holes 40 of defined size from 10 μm to about75 μm, and preferably from about 15 μm to 50 μm, patterned throughout tocoat the at least one surface 52 of the implantable material 56 and istightly adhered to the covered surface 52.

Referring to FIG. 10C, the implantable material 56 covered with the mask48 is then illuminated by the ultraviolet rays. Because TiO₂ issensitive to ultraviolet, the chemical composition in holes 58 isdifferent from the area that is covered by the mask. In contrast to thegeometric features illustrated in FIG. 9C, the geometric features 59 inFIG. 10C is not elevated relative to the surrounding surface of theimplantable material.

Referring to FIG. 10D, after ultraviolet irradiation, the mask isremoved to reveal the surface 52 that surrounds the geometric features59 formed by ultraviolet irradiation. As described above, because theshape of the holes 58 in the mask 48 could be in any of the shapesdescribed for the geometric features including: circle, square,rectangle, triangle, parallel lines and intersecting lines, andcombinations thereof, the geometric features 58 accordingly adopts suchshapes also.

Example 1

Nickel-titanium sheets were heated to oxidize titanium present at thesurface of the sheet. Pattern masks fabricated from machined metal werelaser drilled a pattern of holes having diameters ranging from 15 μm to50 μm, with a single diameter of holes on each pattern mask. A singlepattern mask was placed over a single nickel-titanium sheet and theassembly was exposed to high intensity ultra-violet irradiation. AfterUV irradiation, the irradiated nickel-titanium sheet was placed on afully endothelialized test surface and maintained at 37° C. undersimulated in vivo flow conditions and under static flow conditions.Qualitative observations were periodically made and it was found thatendothelial cells bound to the pattern of UV irradiated affinity domainsand migrated across the nickel-titanium sheet by proliferating acrossthe pattern of affinity domains, eventually fully forming an endotheliumon the nickel-titanium sheet.

What is claimed is:
 1. An implantable biomaterial having at least onesurface, the at least one surface having a first endothelial cellbinding affinity, and a defined pattern of a plurality of functionalfeatures, each of the plurality of functional features having a secondendothelial cell binding affinity that is greater than the firstendothelial cell binding affinity of the at least one surface, whereineach functional feature of the defined pattern of the plurality offunctional features is chemically altered relative to the rest of thesurface of the biomaterial.
 2. The implantable biomaterial according toclaim 1, wherein each of the plurality of functional features furthercomprises a titanium oxide layer.
 3. The implantable biomaterialaccording to claim 1, wherein the biomaterial is made of a material thatcomprises a nickel-titanium alloy.
 4. The implantable biomaterialaccording to claim 1, wherein each of the plurality of functionalfeatures is made of a material that comprises carbon.
 5. The implantablebiomaterial according to claim 1, wherein each of the functionalfeatures comprises a focal adhesion point for affinity binding ofendothelial cells.
 6. The implantable biomaterial according to claim 1,wherein each of the functional features has a width between about 10 μmto about 75 μm.
 7. The implantable biomaterial according to claim 1,further comprising a gap distance between adjacent functional featuresbetween about 10 μm to about 75 μm.
 8. An implantable biomaterial,comprising a nickel-titanium alloy and having at least one surface and aplurality of functional features, each of the plurality of functionalfeatures having a surface energy that is greater than the surface energyof the nickel-titanium alloy.
 9. The implantable biomaterial of claim 8,wherein the plurality of functional features are configured to enhancefocal point adhesion of cells on the implantable biomaterial.
 10. Theimplantable biomaterial according to claim 8, wherein each of theplurality of functional features further comprises a titanium oxidelayer.
 11. The implantable biomaterial according to claim 8, whereineach of the plurality of functional features is made of a material thatcomprises carbon.
 12. The implantable biomaterial according to claim 8,wherein each of the functional features has a width between about 10 μmto about 75 μm.
 13. The implantable biomaterial according to claim 8,further comprising a gap distance between adjacent functional featuresbetween about 10 μm to about 75 μm.